Optical coherence tomography (OCT) is a powerful and sensitive tool for characterization of optical properties and imaging of superficial tissue, as described in the paper by D. Huang, E. A. Swanson, C. P. Lin, J. S. Schuman, W. G. Stinson, W. Chang, M. R. Hee, T. Flotte, K. Gregory, C. A. Puliafito and J. G. Fujimoto, ‘Optical coherence tomography’, published in Science 254, (1991) 1178-1181.
OCT can achieve micrometer depth resolution and allows accurate in-vivo measurement of thickness, area and volume in the tissue. In OCT, the depth dimension is explored by scanning the optical path difference (OPD) between the object path and reference path in an interferometer illuminated by a low coherence source. The maximum interference signal is obtained for OPD=0. In OCT, achievable depth resolution is given by the optical source line-width and not by the numerical aperture of the lens, as is the case in the confocal microscopy. For OPD values larger than the coherence length of the source used, the strength of the interference signal diminishes considerably. This explains the selection in depth of the OCT. Using a superluminescent diode (SLD) OCT depth resolution better than 15 μm is achievable. Employing a larger bandwidth source, 2 μm depth resolution becomes possible, as described in W. Drexler, U. Morgner, R. K. Ghanta, F. X. Kartner, J. S. Schuman, J. G. Fujimoto, “Ultrahigh-resolution ophthalmic optical coherence tomography”, Nature Medicine, Vol. 7, No. 4, 502-507, 2001. OCT is an excellent method for high resolution imaging of superficial tissue, with penetration depths of up to 2-3 mm, depending on the scattering and absorption properties of the tissue.
A reflectivity depth profile called an A-scan is obtained by axial scanning. This means changing the OPD in the interferometer, for instance by moving the reference mirror in the reference arm.
B-scan images, which are analogous to ultrasound B-scan, are generated by collecting many A-scans for different and adjacent transverse positions, a method used in the paper by Huang mentioned above. The lines in the raster correspond to A-scans, i.e. the lines are oriented along the depth coordinate. The transversal scanner (operating along X or Y, or along the radius ρ or the polar angle θ in polar coordinates) advances at a slower pace to build a B-scan image.
Alternatively, a B-scan can be generated by using T-scans. In this case, the transversal scanner produces the fast lines in the image, as described in the papers by A. Gh. Podoleanu, G. M. Dobre, D. J. Webb, D. A. Jackson, “Coherence imaging by use of a Newton rings sampling function,” Opt. Lett. 21, 1789-1791 (1996), by A. Gh. Podoleanu, G. M. Dobre, and D. A. Jackson, “En-face coherence imaging using galvanometer scanner modulation,” Opt. Lett., 23, 147-149 (1998), and by A. Gh. Podoleanu, M. Seeger, G. M. Dobre, D. J. Webb, D. A. Jackson and F. Fitzke “Transversal and longitudinal images from the retina of the living eye using low coherence reflectometry,” J. Biomed Optics, 3, 12-20 (1998). A T-scan can be produced by controlling either the transverse scanner along the X-coordinate, or along the Y-coordinate or along the radius ρ or the polar angle θ with the other transverse and axial scanners fixed. For instance, a T-scan based B-scan is obtained by driving the X-scanner to produce T-scans while the axial scanner advances slower in depth along the Z-coordinate.
A profile of reflectivity obtained while the depth scanning is fixed is called a T-scan. C-scans are made from many T-scans along either of X, Y, ρ or θ coordinates repeated for different values of the other transverse coordinate, Y, X, ρ or θ respectively in the transverse plane. The repetition of T-scans along the other transverse coordinate is performed at a slower rate than that of the T-scans, called the frame rate. In this way, a complete raster is generated. Different transversal slices are collected for different depths Z, either by advancing the optical path difference in the OCT in steps after each complete transverse (XY) or (ρ, θ) scan, or continuously at a much slower speed than the frame rate, as described in the paper by A. Gh. Podoleanu, J. A. Rogers, D. A. Jackson, S. Dunne, “Three dimensional OCT images from retina and skin”, Opt. Express, Vol. 7, No. 9, 292-298, (2000), http://www.opticsexpress.org/abstract.cfm?URI=OPEX-7-9-292.
Typical errors in OCT imaging arise as a result of the specific way in which the image is constructed, i.e. from points of equal OPD, as described in the paper by M. Ohmi, K. Yoden and M. Haruna, “Optical reflection tomography along the geometrical thickness”, Proc. SPIE, Vol. 4251, (2001), 76-80. Such errors are accumulated over the depth in the tissue and lead to deviations of zero OPD points from the position of the focus in confocal microscopy, as described in the paper by R. J. Zawadzki, C. Leisser, R. Leitgeb, M. Pircher, A. F. Fercher, 3D Ophthalmic OCT with a refraction correction algorithm, to be published in Proceed. SPIE, European Conf. Biomedical Optics, 22-25 Jun. 2003, paper 5140-04. The error can exceed the depth resolution achievable with Kerr lens mode-locked lasers, and in some cases, even the resolution achievable with superluminiscent diodes. The lateral errors also can amount to several pixels in the transverse section. These show that correction of images is paramount in order to obtain accurate, interpretable OCT images from the tissue. Diagnosis of glaucoma and macula degeneration relies on the instrument accuracy in determining the retina thickness. There is an increase demand for improving the resolution of the images collected.
Let us consider a curved surface separating two media of different indexes of refraction. The intersection of this surface with the plane y=0 is described by the contour Σ in FIG. 1. For any object point O(x,z), an image point I(h,v) is generated. We orient the axes of the object space and image space along parallel directions. The ray refracted in A, intersects the object point O. The distance in the medium is AO.
Let us consider the medium homogenous of index of refraction n. The frame grabber puts the image point I in the plane (h,v) at a distance from the image point A, equal to the distance AO multiplied by the index of refraction, n. We define two errors, an axial and a lateral error. For instance, in the case of an A-scan, the user expects to collect points from along the line AI in FIG. 1, but in fact the OCT system acquires data along the line AO and places them along the line AI.
When performing a B-scan, if the normal to the surface Σ deviates from the plane of FIG. 1, the B-scan will contain points from the volume outside the plane of FIG. 1.
For C-scanning, the user expects to collect an image from a plane II, perpendicular to the axis OZ. However, due to the curvature of the surface Σ, the coherence gate selects points from inside the object situated on a curved surface, S.
Superposing the origins of the object space and image space in point C, the lateral error El and the axial error Ea are defined as:El=|x−h|  (1a)Ea=|z−v|  (1b)
The axial error includes the elongation of the image in depth due to the index of refraction, n, of the medium or different media intersected by the ray up to the point O.
Optical coherence tomography (OCT) images are collected from the retina with different scanning procedures. The present invention relates to B-scan images, i.e. images containing the optic axis and oriented in depth. First OCT images of the retina have been produced as B-scans, constructed from many A-scans at different transverse positions. An A-scan is a profile of reflectivity in depth.
Development of en-face OCT leads to scanning the retina transversally or angularly. By putting together many T-scans for different depth positions, again a B-scan image of the tissue is obtained.
However, all images so far have been presented as rectangular, i.e. as made of line oriented laterally and axially at 90 degrees. In reality, the eye ball is round.